Adjustment of ventilator pressure-time profile to balance comfort and effectiveness

ABSTRACT

The invention is a ventilator whose servo-controller adjusts the degree of support by adjusting the profile of the pressure waveform as well as the pressure modulation amplitude. As the servo-controller increases the degree of support by increasing the pressure modulation amplitude, it also generates a progressively more square, and therefore efficient, pressure waveform; when the servo-controller decreases the degree of support by decreasing the pressure modulation amplitude, it also generates a progressively more smooth and therefore comfortable pressure waveform.

FIELD OF THE INVENTION

This invention relates to the field of mechanical ventilation, and moreparticularly to machines and methods for providing a patient withventilatory support.

BACKGROUND

Conventional ventilators provide ventilatory support by utilizing anumber of different pressure-time profiles. In its simplest form, aventilator delivers airflow at a fixed rate (or some other fixedfunction of time such as sinusoidally), and the airway pressureincreases passively as a function of the mechanical properties of thepatient's respiratory system. Such a ventilator is in general suitableonly for a paralyzed and sedated patient who cannot change his/herventilation at will. Also, the system is intolerant of leak, so isunsuitable for non-invasive (mask) ventilation.

A bilevel ventilator uses a square pressure-time waveform:

P=P ₀ +A, f>0

P=P₀, otherwise

where P₀ is an end expiratory pressure, chosen to splint the upper andlower airways and alveoli, A is a fixed pressure modulation amplitudechosen to supply a desired degree of support, and f is respiratoryairflow. Here, and throughout what follows, inspiratory flow is definedto be positive, and expiratory flow is defined to be negative. Withbilevel support, the patient can breathe as much or as little as hewishes, by using greater or lesser effort, and the system is somewhatless affected by leak. Some known ventilators, for example, the Servo300 available from Siemens Medical, Iselin, N.J., and the VPAP-ST fromResMed, San Diego Calif., have an adjustment for changing the initialrate of rise of pressure, with the intention of providing a morecomfortable waveform by using a slower rate of rise. In such prior art,the clinician selects a particular waveform, but thereafter the waveformdoes not change, and there is no automatic adjustment of the waveform.

Moving on in complexity, a proportional assist ventilator providespressure equal to an end expiratory pressure P₀ plus a resistance Rmultiplied by respiratory airflow, plus an elastance E multiplied by thetime integral of respiratory airflow:

P=P ₀ +Rf+E∫fdt, f>0

P=P ₀ +Rf, otherwise

(where the integral is from the time of start of the current inspirationto the current moment) in which the resistance R is chosen to unloadsome or all of the resistive work of breathing, and the elastance E ischosen to unload some or all of the elastic work of breathing (that isto say, the Rf term provides a pressure increment to offset some or allof the effort required to get air to flow through the mechanicalpassageways, and the integral term provides some or all of the pressurerequired to overcome the elastic recoil or springiness of the lungs andchest wall). A proportional assist ventilator amplifies patient effort,delivering a natural-feeling waveform, and it is easier for the patientto increase or decrease his ventilation than in the case of bilevelsupport. However, a proportional assist ventilator is disadvantageousfor a patient with abnormal chemoreflexes, as inadequate support isprovided during pathological reductions in effort such as central apneasand hypopneas.

Another approach is to provide a pressure-time profile that iscontinuous function of phase in the respiratory cycle:

P=P ₀ +AΠ(Φ),

where Π(Φ) is a waveform template function, for example, as shown inFIG. 1, and φ is the phase in the respiratory cycle. In FIG. 1, thewaveform template is a raised cosine during the inspiratory part of thecycle, followed by a quasi-exponential decay during the expiratoryportion. This shape will produce a quasi-normal and thereforecomfortable flow-time curve if applied to a passive patient with normallungs.

For example, a servo-ventilator can be constructed by setting thepressure modulation amplitude A to:

A=−G∫(0.5|f|−V _(TGT))dt,

where G is a servo gain (for example, 0.3 cmH₂O per L/min per second),V_(TGT) is a desired target ventilation (e.g., 7.5 L/min), and theintegral is clipped to lie between A_(MIN) and A_(MAX) (for example, 3and 20 cmH₂O) chosen for comfort and safety. A servo-ventilator has theadvantage of guaranteeing a specified ventilation. By setting A_(MIN) tobe non-negative, the patient can at will comfortably breathe more thanthe target ventilation, but in the event of a failure of centralrespiratory drive, the device will guarantee at least a ventilation ofV_(TGT).

Finally, the advantages of using a waveform template can be combinedwith resistive unloading:

5 P=P ₀ +Rf+AΠ(Φ),

where

A=−G∫(0.5|f|−V _(TGT))dt, 0<=A_(MIN)<=A<=A_(MAX)

as before, giving yet more comfort to an awake patient than in the casepreviously considered, yet without losing a guaranteed minimumventilation of V_(TGT).

A disadvantage of the pressure waveform template shown in FIG. 1 is thatit is less efficient than a square wave. That is to say, for any givenamplitude it provides less ventilatory support than a square wave. Thewaveform of FIG. 1 has only half the area of a square wave of the sameamplitude. This can be a problem in patients who require a very highdegree of support, or in the case of mouth leak, where in order toprovide a desired pressure modulation amplitude at the glottic inlet, amuch higher pressure modulation amplitude must be supplied at the mask.The use of pure resistive unloading is similarly inefficient for thesame reason: the area under the pressure-vs-time curve is only half thatof a square wave of the same amplitude. Even the combination of thesmooth waveform template with resistive unloading is less efficient thana square wave of the same amplitude.

It is a general object of our invention to provide a pressure supportventilator that offers the advantages of using a smooth pressurewaveform template while at the same time compensating for itsdisadvantages.

It is another object of our invention to balance comfort andeffectiveness in a ventilator.

SUMMARY OF THE INVENTION

One broad concept implemented by the invention is to change the pressurewaveform in a way that makes an advantageous trade-off between comfortand efficiency, using a more efficient but less comfortable waveformonly when needed.

One aspect of the invention is a ventilator whose servo-controlleradjusts the degree of support by adjusting the profile of the pressurewaveform, preferably while also adjusting the pressure modulationamplitude.

In particular, the servo-controller increases the degree of support byincreasing the pressure modulation amplitude, and also by generating aprogressively more square, and therefore efficient, pressure waveform;the servo-controller decreases the degree of support by decreasing thepressure modulation amplitude, and by generating a progressively moresmooth and therefore comfortable pressure waveform. The changes inamplitude and squareness can be performed sequentially, or partially orcompletely simultaneously.

BRIEF DESCRIPTION OF THE FIGURES

Further objects, features and advantages of the invention will becomeapparent upon consideration of the following detailed description inconjunction with the drawing, in which:

FIG. 1 depicts a smooth and comfortable ventilator waveform templatefunction;

FIG. 2 depicts illustrative apparatus for implementing the method of theinvention; and

FIGS. 3 and 4 show two alternative variable templates for use in theapparatus of FIG. 2 in accordance with the invention, the shape of eachtemplate being a function of the instantaneous difficulty in ventilatingthe patient.

DETAILED DESCRIPTION OF THE INVENTION

Suitable apparatus for implementing the invention is shown in FIG. 2.The apparatus provides breathable gas at controllable positive pressureto a patient's airway. In the drawing, a blower 10 supplies breathablegas to a mask 11 in communication with a patient's airway via a deliverytube 12 and exhausted via an exhaust 13. Airflow at the mask 11 ismeasured using a pneumotachograph 14 and a differential pressuretransducer 15. The mask flow signal f(t) from the transducer 15 is thensampled by a microprocessor 16. Mask pressure is measured at the port 17using a pressure transducer 18. The pressure signal from the transducer18 is then sampled by the microprocessor 16. The microprocessor sends aninstantaneous mask pressure request (i.e., desired mask pressure) signalP(t) to a servo-controller 19, which compares the pressure requestsignal with the actual pressure signal from the transducer 18 to controla fan motor 20. Microprocessor settings can be adjusted via a serialport, not shown.

It is to be understood that the mask could equally be replaced with atracheotomy tube, endotracheal tube, nasal pillows, or other means ofmaking a sealed connection between the air delivery means and thepatient's airway.

The microprocessor accepts the mask airflow and pressure signals, andfrom these signals determines the instantaneous flow through any leakbetween the mask and patient, by any convenient method. For example, theconductance of the leak may be estimated as the instantaneous maskairflow, low-pass filtered with a time constant of 10 seconds, dividedby the similarly low-pass filtered square root of the instantaneous maskpressure, and the instantaneous leakage flow may then be calculated asthe conductance multiplied by the square root of the instantaneous maskpressure. Respiratory airflow is then calculated as the instantaneousmask airflow minus the instantaneous leakage flow.

Throughout the following discussion, the phase in the respiratory cycleΦ is taken as varing between zero and 1 revolution, with zerocorresponding to start of inspiration and 0.5 corresponding to start ofexpiration.

The desired mask pressure is described by the following equations:

P=P ₀ +Rf+AΠ(Φ)

where:

P₀ is a desired end expiratory pressure chosen to splint the upper andlower airways or alveoli, or to reduce cardiac preload or afterload, forexample, 5 cmH₂O;

R may be zero, but is preferably any value less than the patient'sactual airway resistance;

f is respiratory airflow, measured, for example, using apneumotachograph in the mask, and correcting for leak, for example, asdescribed in the commonly owned International Publication referred toabove;

Φ is the phase in the patient's respiratory cycle;

Π(Φ) is a pressure waveform template, initially set to be similar tothat shown in FIG. 1, for example, comprising a raised cosine followedby an exponential decay.

In a very simple form, suitable for a patient who is making nospontaneous efforts, or in whom the spontaneous efforts can be ignored,the phase Φ simply increases linearly with time, modulo 1 revolution. Ina preferred form, the phase φ is calculated, for example, from therespiratory airflow f using fuzzy logic as taught in the commonly ownedInternational Publication No. WO 98/12965 entitled “Assisted Ventilationto Match Patient Respiratory Need,” referred to above.

An example of a smooth and comfortable pressure waveform template Π(Φ)is shown in FIG. 1. This particular waveform consists of a raised cosinefollowed by a quasi-exponential decay. (Unlike a true exponential decay,the waveform of FIG. 1 falls exactly to zero by the end of expiration,so that there is no step change at the start of the next breath.)

The first reason why the waveform of FIG. 1 is more comfortable than atraditional square wave is that the more sudden changes in pressureassociated with a square wave are more intrusive than the smootherchanges in pressure of FIG. 1.

The second reason why the waveform of FIG. 1 is more comfortable,whereas a traditional square wave is less comfortable, relates toprecise synchronization of the delivered pressure to the patient's ownmuscular efforts. The more precise the synchronization, the morecomfortable the waveform. The term R f in the equation

P=P ₀ +Rf+AΠ(Φ)

given above can be adjusted to obviate some or most of the effortrequired to unload resistive work. By a suitable choice of the amplitudeA, and a suitable waveform Π(Φ), the term A Π(Φ) can be adjusted tounload most of the normal or pathological elastic work at the eupneictidal volume, or alternatively at a minimum desired tidal volume,leaving the patient free to breathe deeper if he wishes. The reason forthis is that a eupneically breathing subject's inspiratory flow-timecurve is quasi-sinusoidal, and therefore the elastic component ofeffort, which is proportional to the integral of flow, is a raisedcosine. For this reason, the waveform of FIG. 1 has a raised cosineduring inspiration. During early expiration, a normal subject's musculareffort does not drop instantaneously to zero, but remains active sometime into expiration, decaying gradually, which maintains a high lungvolume longer than would otherwise be the case, thereby helping keep thealveoli inflated, and also providing a smoother movement of the chestwall. For this reason, the waveform of FIG. 1 has a quasi-exponentialdecay during expiration. An additional advantage of the waveform of FIG.1 over a square wave is that small errors in timing of the start ofinspiration produce negligible errors in the delivered pressure, whereaswith a square wave, a timing error causes the delivered pressure to bewrong by the entire amplitude of the waveform. Therefore, the waveformof FIG. 1 will feel better synchronized to the subject's efforts than asquare wave.

Primary interest is in waveform templates which are nondecreasing duringthe inspiratory half of the cycle, nonincreasing during the expiratoryhalf, and with the first derivative defined everywhere except at thetransitions between inspiration and expiration and vice versa. Ofparticular importance are waveform templates which are families offunctions indexed by a single smoothness parameter K, which can forconcreteness take values between zero (least smooth, or most square) and1 (most smooth). The maximum absolute value of the derivative (of thewaveform template with respect to phase) increases as smoothness Kdecreases. Thus in the family of waveform templates shown in FIGS. 3 and4, each waveform is smoother than the waveform immediately to the left.

As the patient's ventilatory requirements increase, the smooth andcomfortable waveform template changes to a progressively more square(and therefore more efficient, but generally less comfortable) waveform.In a preferred form, the pressure waveform template is a function of asmoothness variable K. When K=1.0, the template is smooth as shown inFIG. 1. When K=0.0, the template is a square wave, and intermediatevalues of K generate intermediate waveforms.

FIG. 3 shows one way in which the waveform can vary with K to generateintermediate waveforms. During inspiration, Π(Φ) is a blend between araised cosine and a rising exponential, the time constant of theexponential decreasing with K. During expiration, Π(Φ) is a decayingexponential, also with a time constant that decreases as K increases.Letting, for K>0

u=0.5[1−cos(2πΦ)]

v=a(1−e ^(−5Φ/K))

where

a=1/(1−e ^(−2.5/K))

we define

Π(Φ)=Ku+(1−K)v, Φ<0.5

Π(Φ)=1−a(1−e ^(−5(Φ−0.5)/K)), otherwise.

The equations degenerate to a square wave when K=0. The purpose of theconstant a is to ensure that Π(Φ) approaches zero as Φ approaches 0.5and also as Φ approaches unity.

As K decreases, two things happen to the inspiratory part of the curve:the exponential becomes progressively more like a rising step function,and the exponential contributes progressively more to the template,generating a family of curves intermediate between a raised cosine and asquare wave. Similarly, as K decreases, the exponential in theexpiratory part of the curve becomes more like a descending stepfunction.

FIG. 4 shows another method, in which the inspiratory part of the curveis a raised cosine followed by a straight line:

Π(Φ)=0.5[1−cos(2πΦ/K)], φ<0.5, φ<0.5K

Π(Φ)=1−a(1−e ^(−5(φ−0.5)/K)), φ>0.5

Π(Φ)=1, otherwise

where

a=1/(1−e ^(−2.5/K)).

In this method, with K=1.0, the straight line segment vanishes and theinspiratory curve is a raised cosine. As K decreases, the straight linesegment lengthens and the raised cosine is squashed progressively to theleft. Again, the equations degenerate to a square wave with K=0.0.

In both embodiments, the object is to use K=1.0 when small degrees ofsupport are required, K=0.0 when very large degrees of support arerequired, and intermediate values of K in between.

In a simple form of the invention, K is adjusted in order toservo-control the patient's minute ventilation to equal a chosen target.For example, K may be adjusted using clipped integral control asfollows:

K=G∫(0.5|f|−V _(TGT))dt, 0<=K<=1

where:

G is a gain, for example, 0.01 per L/min per second;

V_(TGT) is the chosen target ventilation, for example, 7.5 L/min;

The reason for dividing the absolute value of the respiratory airflow bytwo is as follows. The target ventilation V_(TGT) is specified with theunits of L/min. Normally, ventilation is calculated as either the entirevolume inspired per minute (inspired minute ventilation), or the entirevolume expired per minute (expired minute ventilation). Equally, it canbe calculated as the average of these two, in which case the averageminute ventilation is half the average of the absolute value of therespiratory airflow over any given minute. More generally, the averageventilation is the average of half the absolute value of the respiratoryairflow over any chosen period of time. Omitting the averaging step, wesee that the instantaneous ventilation is half the absolute value of therespiratory airflow, and the term 0.5|f|−V_(TGT) is the error in theinstantaneous ventilation, and is therefore (on average) a measure ofthe adequacy of ventilation. If the term 0.5|f|−V_(TGT) is on averagepositive, then the subject requires less ventilatory support, andconversely if it is on average negative, then the subject requires moreventilatory support. The clipped integral controller servo-controls thisquantity to be zero on average, and therefore servo-controls theinstantaneous ventilation to on average equal the target ventilation,whereupon the average ventilation also equals the target ventilation.

In this embodiment, if the subject is exceeding the target ventilation,the value of K will increase, yielding progressively smoother, morecomfortable, but less efficient waveforms, until either the actualventilation decreases to equal the target ventilation, or until Kreaches 1.0, which yields the smoothest waveform. Conversely, if thesubject is not achieving the target ventilation, K will decreasegradually, causing the waveform to become more square and moreefficient, until either the target ventilation is achieved, or untilK=0.0, representing a perfectly square waveform. For example, if K=1.0,V_(TGT)=7.5 L/min, G=0.01 per L/min per second, and the subject ceasesall respiratory airflow, K will decrease to zero in 13.3 seconds.

There are two ways of increasing the degree of ventilatory support:using a more square waveform, and increasing the pressure modulationamplitude A. Therefore, in the present invention, both the smoothness Kand the pressure modulation amplitude A may be adjusted, eithersimultaneously or sequentially, in order to achieve synergistically adesired target ventilation.

In a preferred form, a smooth waveform is used preferentially, and asfar as possible the desired target ventilation is achieved by modulatingthe amplitude A, but if this is unsuccessful, then a progressively moresquare waveform is used, by decreasing K. In accordance with this formof the invention, the pressure modulation amplitude A may be adjustedusing a clipped integral controller in order to servo-control minuteventilation to equal a desired target ventilation as follows:

A=−G∫(0.5|f|−V _(TGT))dt, 0<=A_(MIN)<=A<=A_(MAX)

where:

G is a gain, for example, −0.3 cmH₂O per L/min per second;

V_(TGT) is a chosen guaranteed minimum (target) ventilation, forexample, 7.5 L/min;

A_(MIN) is a minimum pressure modulation amplitude, chosen to make thepatient comfortable while awake, for example, 3 cmH₂O; and

A_(MAX) is a maximum pressure modulation amplitude, chosen to besufficient to do all respiratory work, within the constraints oftolerability and safety, for example, 20 cmH₂O.

In the case where the patient's ventilation exceeds the target V_(TGT),the pressure modulation amplitude A will reduce, until either theventilation on average equals V_(TGT) and A lies in the rangeA_(MIN)<A<A_(MAX), or until A reaches A_(MIN). Conversely, in the casewhere A_(MAX) is insufficient to ventilate the patient at V_(TGT). Awill become equal to A_(MAX).

In this preferred form, K is then calculated as a decreasing function ofthe pressure modulation amplitude A. In other words, as the pressuremodulation amplitude A increases with the need for greater ventilatorysupport, K decreases to provide still further support (at the expense ofcomfort). Therefore, the pressure waveform template Π(Φ) becomes afunction of the pressure modulation amplitude A. The invention deliversa comfortable, smooth pressure-vs-phase (and therefore pressure-vs-time)curve, providing the target ventilation V_(TGT) is being achieved with apressure modulation amplitude less than a chosen maximum A_(MAX), butusing a progressively more efficient, and therefore more square,waveform in the case where the target ventilation cannot be achievedusing the chosen maximum.

To this end, the smoothness K may be calculated using clipped integralcontrol using the following pseudocode:

K = 1.0 REPEAT every 20 milliseconds   Calculate A   IF A < A_(MAX)    Increment K by 0.002   ELSE     Decrement K by 0.002   END  Truncate K to lie between 0.0 and 1.0 END

Initially, K=1.0, and the smoothest waveform will be used. In the casewhere the patient is being well ventilated at or above the targetventilation V_(TGT), K will remain at 1.0 and the patient will continueto receive a very smooth and comfortable pressure waveform.

If the patient becomes difficult to ventilate, for example, due tosputum retention, failure of respiratory drive, diaphragm fatigue,failure of accessory muscles of respiration, mouth leak, or a large leakwhich is exceeding the capacity of the blower, K will gradually decreasetowards zero.

The effect is that the actual delivered pressure waveform Π(Φ) changesgradually and continuously between the comfortable and smooth shape inFIG. 1, and a less comfortable but more efficient square wave. In themost severe case with ventilation remaining below V_(TGT) and Aremaining below A_(MAX), K will reach zero in about 10 seconds, and asquare wave will be delivered. In a less severe case, as K decreases andthe waveform becomes progressively more square, and therefore moreefficient at generating ventilation, V_(TGT) will be achieved at anintermediate value of K and therefore at an intermediate waveform shape.

Should the conditions which led to the requirement for a more efficientwaveform subside, the target ventilation V_(TGT) will be met, pressuremodulation amplitude will reduce to below A_(MAX), and K will againincrease, yielding a smoother and more comfortable waveform.

In the example given above, K increases at a maximum rate of 0.1 persecond. Larger rates of change will produce a more rapid increase ineffectiveness of ventilatory support, but are likely to lead toovershoot, with oscillations in the degree of support. Smaller rates ofchange will be stable, but will take longer to re-establish ventilationat V_(TGT).

In the above algorithm, the ventilator attempts to cope with a need forincreased ventilatory support in two discrete stages, first byincreasing the pressure modulation amplitude, while maintaining thesmooth waveform, but only up to a preset maximum amplitude A_(MAX), andthen subsequently by using a progressively more efficient waveform. Inother embodiments it is possible for the two stages to overlap. Forexample, the pseudocode could be changed to:

K = 1.0 REPEAT every 20 milliseconds   Calculate A   Decrement K by0.002(A − 0.75 A_(MAX))   Truncate K to lie between 0.0 and 1.0 END

This algorithm performs identically to the previous algorithm for theextreme cases of a patient who is either very difficult or very easy toventilate, but differs for intermediate cases, because the transitionfrom smooth to square begins earlier, at 75% of A_(MAX). If more than75% of the maximum pressure modulation is being used, K will decrease,and the waveform will become more square. Conversely, if more than 75%of the maximum pressure modulation is being used, K will increase andthe waveform will become more rounded. Thus when increasinglyventilating the patient, it is possible to adjust the trade-off betweenincreasing the pressure modulation amplitude and using a more efficientwaveform.

In some cases, it may be desirable to prevent K from reaching zero. Forexample, keeping 0.1<K<1.0 can produce almost as great an increase inefficiency at low K, but is more comfortable to the patient than acompletely square waveform. This is particularly the case if largeamounts of resistive unloading are used. (This is because a near-squarewaveform template on its own will produce a rapid increase in flow atstart of inspiration, which will then produce yet further increase inpressure due to resistive unloading.)

Alternatively, K can be made to increase quickly at first, and then moreslowly, so that the most square waveform is used only as a last resort,for example, by submitting K to a square root or similar transform. Inother cases, with patients with considerable air hunger and intrinsicPEEP, it may be desirable to limit K to a value less than 1.0, althoughin general it would be preferable to increase the resistive unloading Rand the end expiratory pressure P₀. In the above embodiments, K isrelated to the integral of A (minus a threshold) with respect to time,essentially using an integral controller to determine K, in an attemptto servo-control ventilation to equal or exceed V_(TGT). In otherembodiments, other known controllers such as PID controllers may beused.

Although the invention has been described with reference to particularembodiments, it is to be understood that these embodiments are merelyillustrative of the application of the principles of the invention. Inthe above preferred embodiments, the pressure waveform is a function ofphase in the respiratory cycle φ, calculated as as taught in thecommonly owned International Publication No. WO 98/12965 entitled“Assisted Ventilation to Match Patient Respiratory Need”. However, if itis not desired to synchronize with the patient's spontaneous efforts,phase can be taken as increasing linearly with time at a preset rate,modulo 1 revolution. In this manner, the pressure waveform is a simplefunction of time, and the invention simplifies to modifying the shape ofa fixed pressure-vs-time waveform. Thus, the pressure waveform may be afunction of the phase in the patient's respiratory cycle, or time, or ofboth. Similarly, in the above preferred embodiments, linear resistiveunloading is used, but the invention is applicable in the case of noresistive unloading, and also in the case of nonlinear resistiveunloading. In the preferred embodiments described above, the pressurewaveform template comprises a raised cosine followed by aquasi-exponential decay. However, the precise waveform is not overlycritical. Waveforms with the broad general features of FIG. 1 aresatisfactory, and will generally produce large improvements in comfortand synchronization over a square wave. The waveform may be modified tomore or less precisely include eupneic resistive unloading by settingthe waveform to be more or less closely the shape of the subject'seupneic transdiaphragmatic pressure vs phase curve. Again, in thepreferred embodiments, a fixed non-zero end expiratory pressure is used,yet the invention extrapolates naturally to either zero end expiratorypressure or to automatically adjusted end expiratory pressure.Similarly, some specific examples of how to adjust the shape of thepressure waveform are given, but these are intended only as examples. Inthe examples given, both the inspiratory and expiratory phases of thepressure waveform template increase or decrease their smoothness. In theillustrative embodiments, a single parameter K defines the shape of thepressure waveform template. Also contemplated are embodiments in whichmore than one parameter defines the template. One example is to use oneparameter for inspiration and another for expiration, and vary themindependently. Another is to use one parameter which chiefly affectsearly inspiration and expiration, and another which chiefly affects lateinspiration and expiration. Although the specific implementationdelivers air from a blower, the invention works equally well with air,oxygen, or other breathable gases, and any source of breathable gas atcontrollable pressure may be used. Numerous other modifications may bemade in the illustrative embodiments of the invention and otherarrangements may be devised without departing from the spirit and scopeof the invention.

1-56. (canceled)
 57. A method of controlling a pressure supportventilator comprising the steps of: supplying breathable gas to apatient's airway at a pressure that varies during each respiratory cyclein accordance with an adjustable pressure waveform template having ashape and a variable amplitude, determining adequacy of ventilation ofthe patient, and automatically making a change to the amplitude andshape of said pressure waveform template in accordance with thedetermined adequacy of ventilation, wherein said change produces adecrease in ventilatory support and an increase in comfort when theventilatory support exceeds that required by the patient, and whereinsaid pressure waveform template has both its inspiratory and expiratoryparts defined by a single smoothness parameter whose value isautomatically varied in accordance with the adequacy of patientventilation.
 58. A method in accordance with claim 57 wherein over atleast a portion of each of successive respiratory cycles the amplitudeand shape of the pressure waveform template change in directions thatcause a ventilatory support change in the same direction.
 59. A methodin accordance with claim 58 wherein flow of said breathable gas isservo-controlled to adjust said amplitude so that ventilation equals atarget value, and the shape of said pressure waveform template iscontrolled to be a function of said amplitude.
 60. A method inaccordance with claim 57 wherein flow of said breathable gas isservo-controlled so that ventilation equals a target value, and theshape of said pressure waveform template is controlled to be a functionof flow.